The aim of nuclear-medical imaging is to display physiological and biochemical processes within the body. The patient is in this case given a tracer with a radionuclide, which is distributed in the body and in this case emits radioactive radiation. This radiation is measured using a camera which contains suitable detectors, and the tracer distribution in the body is determined in this way.
In the case of positron emission tomography (PET), positron emitters are used as tracers and emit positrons which decompose in the body into two opposing Gamma quanta. In the case of SPECT (Single Photon Emission Computed Tomography) imaging, in contrast, gamma emitters are used as radionuclides. In both cases, the gamma quanta are trapped, for example, by a matrix of scintillation crystals, in which an impacting photon produces a light flash. This is in turn trapped and amplified by photodetectors, for example by photomultiplier tubes (PMT) or avalanche photodiodes (APD).
Each detector is followed by electronic pre-amplification and filtering of the signals, which are then passed in parallel via a cable harness to the downstream evaluation units (trigger, time and amplitude characterization, energy selection coincidence detector).
Thus, in the known systems, the output signal from each photodetector is carried on a separate signal line. When a PET camera is installed in an MR magnet, a very large number of signal lines must therefore be routed out of the magnet bore in parallel. Since the lines are coupled to the radio-frequency and gradient fields of the MR system, this can lead to signal corruption. Digital processing of the events directly in the magnet will admittedly greatly reduce the amount of information and the number of lines. However, this is virtually impossible in practice because the MR antennas are highly sensitive to interference injected from clock and data signals.
So-called optical multiplexing is already known in order to reduce the number of signal lines in nuclear medicine. In this case, the spatial resolution of the camera is enhanced by way, for example, of a block with a large number of scintillation crystals acting via a light distributor on a group of only four photodetectors, which are fitted at the corners of the block. The light flash which is initiated by a radiation quantum is thus distributed over different photodetectors. A centroid, which corresponds to the location of the absorption of the quantum can be calculated from the signal level of the individual photodetectors. There are a plurality of variants of this method:
In the case of a block detector, one group of four photodetectors is fitted at the corners of a block of, for example, 8×8 crystal segments. The sum of the detector signals indicates the event energy, and the coordinates of the active pixel within the block can be calculated from the difference signals, if necessary after calibration of the position relationship (Anger camera principle).
In the case of a panel detector, there is no optical isolation between the blocks, and the light can be distributed over a plurality of photodetectors located in the vicinity. The intensities result from the position of the active pixel, convolved with a spatial impulse response of the light distributor. One advantage over the block principle is that the sensitivity range of a detector is not artificially restricted to two of four incident quadrants, that is to say all four adjacent blocks are covered. In consequence, only one fourth of the number of detectors required in the case of a block detector are required for a given resolution (for example one detector for 64 pixels).
The number of pixels which can be resolved per photodetector cannot, however be increased indefinitely because of the finite accuracy of the pulse amplitude measurement (photon statistics, detector and amplifier noise).
Methods for combination of the signals from a plurality of detectors onto as few signal lines as possible have already been developed, for example for video cameras. Since these measures do not, however, have to operate in strong magnetic fields, as in the case of a PET camera in an MR magnet, more complicated signal processing methods can be used here. One such method is disclosed, for example, in Kejzlar, Ludek; Fischer, Jan; “Signal Processing by Inherent FIR Filter in the Line CCD Sensor” In: Nové smery v spracovani signálov VI. Liptovský Mikulá{hacek over (s)}, Slowakei, Military Academy, 2002, vol. 2, pp. 215-218.
Frank Gray developed the so-called Gray code as an alternative to normal binary code, see U.S. Pat. No. 2,632,058. The Gray code was used, because it involved only single steps, as early as when finding a solution to the correspondence problem in the calculation of 3D coordinates with the aid of photogrammatic techniques from image data acquired using cameras for reliable inspection of industrial parts, see Gühring, Jens; “Reliable 3D Surface Acquisition, Registration and Validation using Statistical Error Models”. 3DIM 2001, Quebec City, Canada, 2001. Conference Proceedings, pp. 224-231.
One example of a detector arrangement for a PET camera or SPECT camera with photodetectors, scintillation crystals (LSO crystals) and optical fibers is described in Garcia, Ernest V.; Faber, Tracey L.; Galt, James R. et al.: Advances in Nuclear Emission PET and SPECT Imaging. IEEE Engineering in Medicine and Biology Magazine, Vol. 19, No. 5, September-October 2000, pp. 21-33.